This invention relates to radiation imaging methods and systems and, more particularly, to such imaging systems utilizing arrays of scintillation detectors in combination with fluorescent optical converters.
Radiologic and radionuclide imaging have, within the past decade, entered a period of rapid growth and development. Much of this activity is the result of the introduction of computerized tomography and new imaging instruments and methods.
One of the earlier imaging systems, the Anger scintillation camera (U.S. Pat. No. 3,011,057), was introduced in the 1950's and various improvements have been made upon this basic instrument. With the introduction of relatively cheap and high-speed computers, further processing of scintillation camera images to enhance their diagnostic utility has become routine.
The conventional Anger camera consists of an array of photomultipliers coupled via a flat light pipe to a large, thin, flat scintillation crystal which may be, for example, thallium-activated sodium iodide (NAI(Tl)). A collimator is placed in front of the scintillator. Radiation passing through the collimator impinges upon the scintillator and is converted into optical radiation whose spectrum extends from the ultraviolet to the blue portion of the visible spectrum. This optical radiation then passes via a fiber optic coupling into photomultipler detectors where it is converted into electrical signals whose summed pulse-height voltage is proportional to the energy of the absorbed ionizing radiation photon. The position of the scintillation event is estimated by mixing the individual pulse-height amplitudes from the photomultipliers.
Any detector used for radionuclide imaging must be capable of meeting certain requirements which are:
(1) High spatial resolution in determining the position of the scintillation event;
(2) Sufficient energy resolution to identify and reject gamma-photons that have been scattered in the source or in the collimator (source scatter rejection) as well as within the detector itself (detector scatter);
(3) A reasonable count rate and low dead time;
(4) A reasonable gamma-ray absorption efficiency for the detector (minimum signal loss).
In addition to the above, there are a number of attributes that would be desirable in an imaging system but are often not obtainable which include:
(1) The capability of rejecting gamma-photons that are scattered within the scintillator and subsequently reabsorbed in some other portion of the scintillator;
(2) The capability of using thick scintillators to increase incident gamma-photon absorption rates without significant deterioration in spatial resolution;
(3) The ability of providing digital information for computer storage and processing;
(4) An immunity from drift and instability over time, and inhomogeneity across the face of the detector.
These desirable qualities, along with lower cost, simplified operation, and higher efficiency (in the sense of image quality for a given quantity of radionuclide administered to the patient), have been sought through the use of several alternative imaging schemes. One approach is embodied by the autofluoroscope. This device uses light pipes to couple individual scintillators to an array of photomultipliers and includes several hundred individual scintillators. Similar devices have been constructed as planar arrays for scintigraphy or ring cameras for position emission tomography imaging.
The autofluoroscope and similar devices suffer from several problems. One is that the transmission of light from the scintillator, through the light pipes, to the photomultipliers is inefficient and nonuniform. This results in poor energy resolution and poor source scatter rejection. In addition, such devices are exceedingly complex because each scintillator element must be individually and reproducibly coupled to a light pipe and a light conducting "spatula" resulting in 315 light conduits for an array of 294 scintillators. The number of photomultipliers required to address even a moderate size system of this type remains large.